Intraluminal device with improved flexibility and durability

ABSTRACT

In accordance with the present invention, there is provided a stent for insertion into a vessel of a patient. The stent is a tubular member having front and back open ends and a longitudinal axis extending there between. The tubular member has a first smaller diameter for insertion into a patient and navigation through the vessels, and a second larger diameter for deployment into the target area of a vessel. The tubular member is made from a plurality of adjacent hoops extending between the front and back ends. The hoops include a plurality of longitudinal struts and a plurality of loops connecting adjacent struts. The stent further includes a plurality of bridges having loop to bridge connections which connect adjacent hoops to one another. The bridge to loop connection points are separated angularly with respect to the longitudinal axis. The bridges have one end attached to a loop, another end attached to a loop on an adjacent hoop. The connection point between the bridge and the hoops will have a repeating pattern over a plurality of strut apices such that the benefits of a decreased number of bridges is realized while simultaneously avoiding the creation of overly unconstrained hoops. It is preferred that the ratio of total number of circumferentially aligned loops to the number of loops spanned by a particular bridge be a whole number.

CROSS REFERENCE TO RELATED APPLICATION

This application claims the benefit of U.S. Provisional Application Ser.No. 61/256,633 filed Oct. 30, 2009, which is incorporated by referenceherein.

FIELD OF THE INVENTION

The present invention relates to an expandable intraluminal grafts(“stents”) for use within a body passageway or duct which areparticularly useful for repairing blood vessels narrowed or occluded bydisease. The present invention relates even further to such stents whichare self-expanding and made from a superelastic material such asNitinol. The present invention also relates to delivery systems for suchstents.

BACKGROUND OF THE INVENTION

Percutaneous transluminal coronary angioplasty (PTCA) is a therapeuticmedical procedure used to increase blood flow through the coronaryartery and can often be used as an alternative to coronary by-passsurgery. In this procedure, the angioplasty balloon is inflated withinthe stenosed vessel, or body passageway, in order to shear and disruptthe wall components of the vessel to obtain an enlarged lumen. Withrespect to arterial stenosed lesions, the relatively incompressibleplaque remains unaltered, while the more elastic medial and adventitiallayers of the body passageway stretch around the plaque. This processproduces dissection, or a splitting and tearing, of the body passagewaywall layers, wherein the intima, or internal surface of the artery orbody passageway, suffers fissuring. This dissection forms a “flap” ofunderlying tissue which may reduce the blood flow through the lumen, orblock the lumen.

Typically, the distending intraluminal pressure within the bodypassageway can hold the disrupted layer, or flap, in place. If theintimal flap created by the balloon dilation procedure is not maintainedin place against the expanded intima, the intimal flap can fold downinto the lumen and close off the lumen, or may even become detached andenter the body passageway. When the intimal flap closes off the bodypassageway, immediate surgery is necessary to correct this problem.

Recently, transluminal prostheses have been widely used in the medicalarts for implantation in blood vessels, biliary ducts, or other similarorgans of the living body. These prostheses are commonly known as stentsand are used to maintain, open, or dilate tubular structures. An exampleof a commonly used stent is given in U.S. Pat. No. 4,733,665 filed byPalmaz on Nov. 7, 1985, which is hereby incorporated herein byreference. Such stents are often referred to as balloon expandablestents. Typically the stent is made from a solid tube of stainlesssteel. Thereafter, a series of cuts are made in the wall of the stent.The stent has a first smaller diameter which permits the stent to bedelivered through the human vasculature by being crimped onto a ballooncatheter. The stent also has a second, expanded diameter, upon theapplication, by the balloon catheter, from the interior of the tubularshaped member of a radially, outwardly extending.

However, such stents are often impractical for use in some vessels suchas the carotid artery. The carotid artery is easily accessible from theexterior of the human body, and is often visible by looking at onesneck. A patient having a balloon expandable stent made from stainlesssteel or the like, placed in their carotid artery might be susceptibleto sever injury through day to day activity. A sufficient force placedon the patient's neck, such as by falling, could cause the stent tocollapse, resulting in injury to the patient. In order to prevent this,self expanding stents have been proposed for use in such vessels. Selfexpanding stents act like springs and will recover to their expanded orimplanted configuration after being crushed.

One type of self-expanding stent is disclosed in U.S. Pat. No.4,665,771, which stent has a radially and axially flexible, elastictubular body with a predetermined diameter that is variable under axialmovement of ends of the body relative to each other and which iscomposed of a plurality of individually rigid but flexible and elasticthread elements defining a radially self-expanding helix. This type ofstent is known in the art as a “braided stent” and is so designatedherein. Placement of such stents in a body vessel can be achieved by adevice which comprise an outer catheter for holding the stent at itsdistal end, and an inner piston which pushes the stent forward once itis in position.

However, braided stents have many disadvantages. They typically do nothave the necessary radial strength to effectively hold open a diseasedvessel. In addition, the plurality of wires or fibers used to make suchstents could become dangerous if separated from the body of the stent,where it could pierce through the vessel. Therefore, there has been adesire to have a self-expanding stent, which is cut from a tube ofmetal, which is the common manufacturing method for many commerciallyavailable balloon expandable stents. In order to manufacture aself-expanding stent cut from a tube, the alloy used would preferably besuperelastic or psuedoelastic characteristics at body temperature, sothat it is crush recoverable.

The prior art makes reference to the use of alloys such as Nitinol(Ni—Ti alloy) which have shape memory and/or superelasticcharacteristics in medical devices which are designed to be insertedinto a patient's body. The shape memory characteristics allow thedevices to be deformed to facilitate their insertion into a body lumenor cavity and then be heated within the body so that the device returnsto its original shape. Superelastic characteristics on the other handgenerally allow the metal to be deformed and restrained in the deformedcondition to facilitate the insertion of the medical device containingthe metal into a patient's body, with such deformation causing the phasetransformation. Once within the body lumen the restraint on thesuperelastic member can be removed, thereby reducing the stress thereinso that the superelastic member can return to its original un-deformedshape by the transformation back to the original phase.

Alloys having shape memory/superelastic characteristics generally haveat least two phases. These phases are a martensite phase, which has arelatively low tensile strength and which is stable at relatively lowtemperatures, and an austenite phase, which has a relatively hightensile strength and which is stable at temperatures higher than themartensite phase.

Shape memory characteristics are imparted to the alloy by heating themetal at a temperature above which the transformation from themartensite phase to the austenite phase is complete, i.e. a temperatureabove which the austenite phase is stable (the Af temperature). Theshape of the metal during this heat treatment is the shape “remembered”.The heat treated metal is cooled to a temperature at which themartensite phase is stable, causing the austenite phase to transform tothe martensite phase. The metal in the martensite phase is thenplastically deformed, e.g. to facilitate the entry thereof into apatient's body. Subsequent heating of the deformed martensite phase to atemperature above the martensite to austenite transformation temperaturecauses the deformed martensite phase to transform to the austenite phaseand during this phase transformation the metal reverts back to itsoriginal shape if unrestrained. If restrained, the metal will remainmartensitic until the restraint is removed.

Methods of using the shape memory characteristics of these alloys inmedical devices intended to be placed within a patient's body presentoperational difficulties. For example, with shape memory alloys having astable martensite temperature below body temperature, it is frequentlydifficult to maintain the temperature of the medical device containingsuch an alloy sufficiently below body temperature to prevent thetransformation of the martensite phase to the austenite phase when thedevice was being inserted into a patient's body. With intravasculardevices formed of shape memory alloys having martensite-to-austenitetransformation temperatures well above body temperature, the devices canbe introduced into a patient's body with little or no problem, but theymust be heated to the martensite-to-austenite transformation temperaturewhich is frequently high enough to cause tissue damage and very highlevels of pain.

When stress is applied to a specimen of a metal such as Nitinolexhibiting superelastic characteristics at a temperature above which theaustenite is stable (i.e. the temperature at which the transformation ofmartensite phase to the austenite phase is complete), the specimendeforms elastically until it reaches a particular stress level where thealloy then undergoes a stress-induced phase transformation from theaustenite phase to the martensite phase. As the phase transformationproceeds, the alloy undergoes significant increases in strain but withlittle or no corresponding increases in stress. The strain increaseswhile the stress remains essentially constant until the transformationof the austenite phase to the martensite phase is complete. Thereafter,further increase in stress are necessary to cause further deformation.The martensitic metal first deforms elastically upon the application ofadditional stress and then plastically with permanent residualdeformation.

If the load on the specimen is removed before any permanent deformationhas occurred, the martensitic specimen will elastically recover andtransform back to the austenite phase. The reduction in stress firstcauses a decrease in strain. As stress reduction reaches the level atwhich the martensite phase transforms back into the austenite phase, thestress level in the specimen will remain essentially constant (butsubstantially less than the constant stress level at which the austenitetransforms to the martensite) until the transformation back to theaustenite phase is complete, i.e. there is significant recovery instrain with only negligible corresponding stress reduction. After thetransformation back to austenite is complete, further stress reductionresults in elastic strain reduction. This ability to incur significantstrain at relatively constant stress upon the application of a load andto recover from the deformation upon the removal of the load is commonlyreferred to as superelasticity or pseudoelasticity. It is this propertyof the material which makes it useful in manufacturing tube cutself-expanding stents. The prior art makes reference to the use of metalalloys having superelastic characteristics in medical devices which areintended to be inserted or otherwise used within a patient's body. Seefor example, U.S. Pat. No. 4,665,905 (Jervis) and U.S. Pat. No.4,925,445 (Sakamoto et al.).

However, the prior art has yet to disclose any suitable tube cut selfexpanding stents. In addition, many of the prior art stents lacked thenecessary rigidity or hoop strength to keep the body vessel open. Inaddition, many of the prior art stents have large openings at theirexpanded diameter. The smaller the openings are on an expanded stent,the more plaque or other deposits it can trap between the stent and thevessel wall. Trapping these deposits is important to the continuinghealth of the patient in that it helps prevent strokes as well as helpsprevents restenosis of the vessel it is implanted into. The presentinvention provides for a selfexpanding tube cut stent which overcomesmany of the disadvantages associated with the prior art stents.

SUMMARY OF THE INVENTION

In accordance with the present invention, there is provided a stent forinsertion into a vessel of a patient. The stent is a tubular memberhaving front and back open ends and a longitudinal axis extendingtherebetween. The tubular member has a first smaller diameter forinsertion into a patient and navigation through the vessels, and asecond larger diameter for deployment into the target area of a vessel.The tubular member is made from a plurality of adjacent hoops extendingbetween the front and back ends. The hoops include a plurality oflongitudinal struts and a plurality of loops connecting adjacent struts.The stent further includes a plurality of bridges having loop to bridgeconnections which connect adjacent hoops to one another. The bridge toloop connection points are separated angularly with respect to thelongitudinal axis. The bridges have one end attached to a loop, anotherend attached to a loop on an adjacent hoop. The bridges have anon-linear curved profile between their bridge to loop connectionpoints.

BRIEF DESCRIPTION OF DRAWINGS

The foregoing and other aspects of the present invention will best beappreciated with reference to the detailed description of the inventionin conjunction with the accompanying drawings, wherein:

FIG. 1 is a simplified partial cross-sectional view of a stent deliveryapparatus having a stent loaded therein, which can be used with a stentmade in accordance with the present invention.

FIG. 2 is a view similar to that of FIG. 1 but showing an enlarged viewof the distal end of the apparatus.

FIG. 3 is a perspective view of a stent made in accordance with thepresent invention, showing the stent in its compressed state.

FIG. 4 is a sectional, flat view of the stent shown in FIG. 1.

FIG. 4A is an enlarged view of section of the stent shown in FIG. 4.

FIG. 5 is a perspective view of the stent shown in FIG. 1 but showing itin its expanded state.

FIG. 6 is an enlarged sectional view of the stent shown in FIG. 5.

FIG. 7A is a view similar to that of FIG. 4 but showing an alternativeembodiment of the present invention.

FIG. 7B is a view similar to that of FIG. 4 but showing an alternativeembodiment of the present invention.

FIG. 7C is a view similar to that of FIG. 4 but showing an alternativeembodiment of the present invention.

FIG. 7D is a view similar to that of FIG. 4 but showing an alternativeembodiment of the present invention.

FIG. 7E is a view similar to that of FIG. 4 but showing an alternativeembodiment of the present invention.

FIG. 7F is a view similar to that of FIG. 4 but showing an alternativeembodiment of the present invention.

FIG. 8A is an enlarged view of a bridge member according to oneembodiment of the present invention.

FIG. 8B is an enlarged view of a bridge member according to oneembodiment of the present invention.

DETAILED DESCRIPTION OF THE INVENTION

Referring now to the figures wherein like numerals indicate the sameelement throughout the views, there is shown in FIGS. 3 and 4, a stent50 made in accordance with the present invention. FIGS. 3 and 4 showstent 50 in its un-expanded or compressed state. Stent 50 is preferablymade from a superelastic alloy such as Nitinol. Most preferably, stent50 is made from an alloy comprising from about 50.5% (as used hereinthese percentages refer to atomic percentages) Ni to about 60% Ni, andmost preferably about 55% Ni, with the remainder of the alloy Ti.Preferably, the stent is such that it is superelastic at bodytemperature, and preferably has an Af in the range from about 24.degree.C. to about 37.degree. C. The superelastic design of the stent makes itcrush recoverable which, as discussed above, can be used as a stent orframe for any number of vascular devices for different applications.

Stent 50 is a tubular member having front and back open ends 81 and 82and a longitudinal axis 83 extending therebetween. The tubular memberhas a first smaller diameter, FIGS. 3 and 4, for insertion into apatient and navigation through the vessels, and a second largerdiameter, FIGS. 5 and 6, for deployment into the target area of avessel. The tubular ember is made from a plurality of adjacent hoops 52,FIG. 4A showing hoops 52(a)-52(b), extending between the front and backends 81 and 82. The hoops 52 include a plurality of longitudinal struts60 and a plurality of loops 62 connecting adjacent struts, whereinadjacent struts are connected at opposite ends so as to form a series ofpeaks 78 and valleys 80 in a substantially S or Z shape pattern. Theloops 62 are curved substantially semi-circular and symmetrical sectionshaving centers 64 and a substantially constant radius of curvature inthe crimped configuration illustrated in FIG. 4A. The peak 78 and valley80 are defined as the apex along the outside and inside curve,respectively, of loop member 62.

Stent 50 further includes a plurality of bridges 70 which connectadjacent hoops 52 which can best be described by referring to FIG. 4.Each bridge has two ends 56 and 58. The bridges have one end attached toone strut and/or loop, another end attached to a strut and/or loop on anadjacent hoop. In one embodiment, bridges 70 connect adjacent strutstogether at bridge to loop connection points 72 and 74. For example, end56 is connected to loop 64(a) at bridge to loop connection point 72, andend 58 is connected to loop 64(b) at bridge to loop connection point 74.Each bridge to loop connection point has a center 76. The bridge to loopconnection points are separated angularly with respect to thelongitudinal axis. That is the connection points are not immediatelyopposite each other. One could not draw a straight line between theconnection points, wherein such line would be parallel to thelongitudinal axis of the stent.

The above described geometry helps to better distribute strainthroughout the stent, prevents metal to metal contact when the stent isbent, and minimizes the opening size between the features, struts loopsand bridges. The number of and nature of the design of the struts, loopsand bridges are important factors when determining the workingproperties and fatigue life properties of the stent. Preferably, eachhoop has between 24 and 36 or more struts. Preferably the stent has aratio of number of struts per hoop to strut length L (in inches) whichis greater than 200. The length of a strut is measured in its compressedstate parallel to the longitudinal axis 83 of the stent.

As seen from FIGS. 4 and 5, the geometry of the stent changes quitesignificantly as a stent is deployed from its un-expanded state to itsexpanded state. As a stent undergoes diametric change, the strut angleand strain levels in the loops and bridges are affected. Preferably, allof the stent features will strain in a predictable manor so that thestent is reliable and uniform in strength. In addition, it is preferableto minimize the maximum strain experienced by struts loops and bridges,since Nitinol properties are more generally limited by strain ratherthan by stress as most materials are. As will be discussed in greaterdetail below, the stent sits in the delivery system in its un-expandedstate as shown in FIG. 4. As the stent is deployed, it is allowed toexpand towards its expanded state, as shown in FIG. 5, which preferablyhas a diameter which is the same or larger than the diameter of thetarget vessel. Nitinol stents made from wire deploy in much the samemanor and are dependent upon the same design constraints as laser cutstents. Stainless steel stents deploy similarly in terms of geometricchanges as they are assisted with forces from balloons or other devices.

In trying to minimize the maximum strain experienced by features, thepresent invention utilizes structural geometry's which distribute strainto areas of the stent which are less susceptible to failure than others.For example, one of the most vulnerable areas of the stent is the insideradius of the connecting loops. The connecting loops undergo the mostdeformation of all the stent features. The inside radius of the loopwould normally be the area with the highest level of strain on thestent. This area is also critical in that it is usually the smallestradius on the stent. Stress concentrations are generally controlled orminimized by maintaining the largest radii possible. Similarly, we wantto minimize local strain concentrations on the bridge and bridgeconnection points. One way to accomplish this is to utilize the largestpossible radii while maintaining feature widths which are consistentwith applied forces. Another consideration is to minimize the maximumopen area of the stent. Efficient utilization of the original tube fromwhich the stent is cut increases stent strength and its ability to trapembolic material.

Many of these objectives have been accomplished by a preferredembodiment of the present invention, shown in FIGS. 3, 4 and 7A-7F. Asseen from these figures, the most compact designs which maintain thelargest radii at the loop to bridge connections are non-symmetric withrespect to the centerline of the strut connecting loop. That is, loop tobridge connection point centers 76 are offset from the center 64 of theloops 62 to which they are attached. The feature is particularlyadvantageous for stents having large expansion ratios, which in turnrequires them to have extreme bending requirements where large elasticstrains are required. Nitinol can withstand extremely large amounts ofelastic strain deformation, so the above features are well suited tostents made from this alloy. This feature allows for maximum utilizationof Ni—Ti or other material capabilities to enhance radial strength,improve stent strength uniformity, improves fatigue life by minimizinglocal strain levels, allows for smaller open areas which enhanceentrapment of embolic material, and improves stent apposition inirregular vessel wall shapes and curves.

As seen in FIG. 4A, stent 50 has strut connecting loops 62 having awidth W4, as measured at the center 64 parallel to axis 83, which aregreater than the strut widths W2, as measured perpendicular to axis 83itself In fact it is preferable that the thickness of the loops vary sothat they are thickest near their centers This increases straindeformation at the strut and reduces the maximum strain levels at theextreme radii of the loop. This reduces the risk of stent failure andallows us to maximize radial strength properties. The feature isparticularly advantageous for stents having large expansion ratios,which in turn requires them to have extreme bending requirements wherelarge elastic strains are required. Nitinol can withstand extremelylarge amounts of elastic strain deformation, so the above features arewell suited to stents made from this alloy. This feature allows formaximum utilization of Ni—Ti or other material capabilities to enhanceradial strength, improve stent strength uniformity, improves fatiguelife by minimizing local strain levels, allows for smaller open areaswhich enhance entrapment of embolic material, and improves stentapposition in irregular vessel wall shapes and curves.

As mentioned above bridge geometry changes as a stent is deployed fromits compressed state to its expanded state and vice-versa. As a stentundergoes diametric change, strut angle and loop strain is affected.Since the bridges are connected to either the loops, struts or both,they are affected. Twisting of one end of the stent with respect to theother, while loaded in the stent delivery system, should be avoided.Local torque delivered to the bridge ends displaces the bridge geometry.If the bridge design is duplicated around the stent perimeter, thisdisplacement causes rotational shifting of the two loops being connectedby the bridges. If the bridge design is duplicated throughout the stent,as in the present invention, this shift will occur down the length ofthe stent. This is a cumulative effect as one considers rotation of oneend with respect to the other upon deployment. A stent delivery system,such as the one described below, will deploy the distal end first, andthen allow the proximal end to expand. It would be undesirable to allowthe distal end to anchor into the vessel wall while holding the stentfixed in rotation, then release the proximal end. This could cause thestent to twist or whip in rotation to equilibrium after it is at leastpartially deployed within the vessel. Such whipping action could causedamage to the vessel.

However, one embodiment of the present invention, as shown in FIGS. 3and 4, reduces the chance of such events from happening when deployingthe stent. By mirroring the bridge geometry longitudinally down thestent, the rotational shift of the Z-sections can be made to alternateand will minimize large rotational changes between any two points on agiven stent during deployment or constraint. That is the bridgesconnecting loop 52(b) to loop 52(c) are angled upwardly from left toright, while the bridges connecting loop 52(c) to loop 52(d) are angleddownwardly from left to right. This alternating pattern is repeated downthe length of the stent. This alternating pattern of bridge slopesimproves the torsional characteristics of the stent so as to minimizeany twisting or rotation of the stent with respect to any two hoops.This alternating bridge slope is particularly beneficial if the stentstarts to twist in vivo. As the stent twists, the diameter of the stentwill change. Alternating bridge slopes tend to minimize this effect. Thediameter of a stent having bridges which are all sloped in the samedirection will tend grow if twisted in one direction and shrink iftwisted in the other direction. With alternating bridge slopes thiseffect is minimized and localized.

The feature is particularly advantageous for stents having largeexpansion ratios, which in turn requires them to have extreme bendingrequirements where large elastic strains are required. Nitinol canwithstand extremely large amounts of elastic strain deformation, so theabove features are well suited to stents made from this alloy. Thisfeature allows for maximum utilization of Ni—Ti or other materialcapabilities to enhance radial strength, improve stent strengthuniformity, improves fatigue life by minimizing local strain levels,allows for smaller open areas which enhance entrapment of embolicmaterial, and improves stent apposition in irregular vessel wall shapesand curves.

Preferably, stents are laser cut from small diameter tubing. For priorart stents, this manufacturing process lead to designs with geometricfeatures, such as struts, loops and bridges, having axial widths W2, W4and W3 (respectively) which are larger than the tube wall thickness T(shown in FIG. 5). When the stent is compressed, most of the bendingoccurs in the plane that is created if one were to cut longitudinallydown the stent and flatten it out. However, for the individual bridges,loops and struts, which have widths greater than their thickness, theyhave a greater resistance to this in-plane bending than they do to outof plane bending. Because of this, the bridges and struts tend to twist,so that the stent as a whole can bend more easily. This twisting is abuckling condition which is unpredictable and can cause potentially highstrain.

However, this problem has been solved in the preferred embodiment of thepresent invention, shown in FIGS. 3 and 4. As seen from these figures,the widths of the struts, hoops and bridges are equal to or less thanthe wall thickness of the tube. Therefore, substantially all bendingand, therefore, all strains are “out of plane”. This minimizes twistingof the stent which minimizes or eliminates buckling and unpredictablestrain conditions. The feature is particularly advantageous for stentshaving large expansion ratios, which in turn requires them to haveextreme bending requirements where large elastic strains are required.Nitinol can withstand extremely large amounts of elastic straindeformation, so the above features are well suited to stents made fromthis alloy. This feature allows for maximum utilization of Ni—Ti orother material capabilities to enhance radial strength, improve stentstrength uniformity, improves fatigue life by minimizing local strainlevels, allows for smaller open areas which enhance entrapment ofembolic material, and improves stent apposition in irregular vessel wallshapes and curves.

A number of approaches have been used to add flexibility and durabilityto basic stent design for indications having loading modalitiesinvolving dynamic bending, torsion, and axial extension/compression.Prior concepts, which reduced the number of bridges specified in thestandard stent designs have shown excellent flexibility, but have axialinstability, resulting in problematic catheterization and deploymentcharacteristics. The present invention is directed towards maintainingthe advantages associated with fewer bridges between hoops whileproviding a structural element that leaves enough axial constraintbetween hoops to provide stability during the catheterization anddeployment processes. By lengthening the bridge relative to the lengthof the struts forming the individual hoops, the bridge providesnecessary constraint between adjacent hoops while simultaneouslyproviding the capability to absorb some of the deformation associatedwith torsion, bending and axial extension/compression. In addition, thewidth of the bridge along its length may be “tuned” to maximize thebalance between flexibility and cyclic deformation such that durabilitymay be optimized while still tolerating the presence of non-radialforces within the stent structure.

One alternate embodiment of the present invention that adds thisflexibility while still maintain axial constraint and stability is shownin FIGS. 7A through 7F. FIGS. 7A through 7F show stent 150 which issimilar to stent 50 shown in the previous drawings. Stent 150 is madefrom a plurality of adjacent hoops 152, FIGS. 7A through 7F show hoops152(a)-152(d). The hoops 152 include a plurality of longitudinal struts160 and a plurality of loops 162 connecting adjacent struts, whereincircumferentially adjacent struts are connected at opposite ends so asto form series of peaks or apices and valleys in an S or Z shapepattern. Stent 150 further includes a plurality of bridges 170 whichconnect adjacent hoops 152 at bridge to loop connection points. As seenfrom FIGS. 7A-F and FIGS. 8A-B, bridges 170 incorporate elongated linearstrut sections 175 connected on each end to a first end of a curvedbridge loop members 180. The second end of the curved loop member 180 isconnected to the adjacent hoop 152 at the bridge to loop connectionpoint. In a one embodiment, the curved loop member 180 is attached tothe loop 162 of the adjacent hoop 152.

In a preferred embodiment, the ratio of the circumference of the hoop152 to the length of the elongated linear strut member 175 is less than5.

Each bridge 170 is sized to span a plurality of loops 162 between theconnection points on adjacent hoops 152. The configuration providesadditional structural stability in the open area between adjacent hoops152. The elongated bridge members 170 may be designed to approximate themechanical behavior of a helical spring coil. The result is a stent 150that has repeating hoop sections 152 for radial strength, and bridgesections 170 that provide needed flexibility under bending andaxial/torsional loading conditions.

Individual hoops 152 tend to be unstable when deformed by typicalcatheterization and deployment forces, so the connection point betweenthe bridge 170 and the hoops 152 are located to avoid creating largeareas of axially unconstrained strut apices. In a preferred embodiment,the ratios of the hoop circumference to the distance between theadjacent hoops is between 20:1 and 50:1, and preferably about 25:1.

In a preferred embodiment the connection point between the bridge 170and the hoops 152 will have a repeating pattern over a plurality ofloops 162 such that the benefits of a decreased number of bridges 170 isrealized while simultaneously avoiding the creation of overlyunconstrained hoops 152. It is preferred that the ratio of total numberof loops 162 per side (proximal or distal) of the hoop 152 to the numberof loops 162 (per side) having connection regions (also spanned by aparticular bridge 170) for the given hoop 152 be a whole number. Forexample, FIGS. 7A-F depict a stent having 16 loops 162 per side of thehoop 152. A preferred embodiment would have 8, 4 or 2 connection regions(i.e. bridge 170 would span 2, 4, or 8 loops 162) on the given side andmaintain symmetry. The selected ratio should be chosen to maximizeflexibility and structural stability. FIG. 7A depicts a stent having 16loops 162 per side of the hoop 152 with a total of 8 bridges (8connection regions per side), each spanning 2 loops 162. FIG. 7B depictsa stent having 16 loops 162 per side of hoop 152 with a total of 4bridges (4 connection regions per side), each spanning 4 loops 162. FIG.7C depicts a stent having 16 loops 162 per side of hoop 152 with a totalof 2 bridges (2 connection regions per side), each spanning 8 loops 162.FIGS. 7A-C depict adjacent hoops 152 to be in axial alignment. That iseach loop 162 on each hoop 152 is in the same orientation relative tothe longitudinal axis. However, the loops 162 on adjacent hoops 152 mayrotationally offset, i.e. not in axial alignment to provide longerstruts and added flexibility. The stents 150 depicted in FIGS. 7D-Fillustrate hoop sections 152 that are rotationally offset from theadjacent hoop section 152. In particular, this rotational offset isequal to a 180 degree phase shift, yielding adjacent hoops 152 that area mirror image of one another.

The width of the elongated linear strut section 175 may vary along itslength, preferably being symmetrical about its center to avoidnon-uniform deflection characteristics. The connection point between thebridge 170 and hoops 152 is likely to form a natural hinge point underdeflection. In a preferred embodiment, the bridge 170 width at theconnection point to the hoop 152 will be optimized such that fatiguedurability is reasonably maintained. To achieve this optimization, thebridge 170 width at the connection point will be wider than other pointsalong the length of the bridge 170. The bridge 170 shape and width maybe further optimized to reduce out-of-plane forces that develop as aresult of torsion, e.g. a bridge 170 may have its narrowest point at itscenter (relative to its length) to reduce the amount of torsionaldistortion transmitted between hoops 152. FIG. 8B illustrates a bridge170 having a tapered elongated linear strut section 175 with thenarrowest point at the center point along its length.

As mentioned above, it is preferred that the stent of the presentinvention be made from a superelastic alloy and most preferably made ofan alloy material having greater than 50.5 atomic % Nickel and thebalance titanium. Greater than 50.5 atomic % Nickel allows for an alloyin which the temperature at which the martensite phase transformscompletely to the austenite phase (the Af temperature) is below humanbody temperature and preferably is about 24.degree. C. to about37.degree. C. so that austenite is the only stable phase at bodytemperature.

In manufacturing the Nitinol stent, the material is first in the form ofa tube. Nitinol tubing is commercially available from a number ofsuppliers including Nitinol Devices and Components, Fremont Calif. Thetubular member is then loaded into a machine which will cut thepredetermined pattern of the stent, which was discussed above and isshown in the figures, into the tube. Machines for cutting patterns intubular devices to make stents or the like are well known to those ofordinary skill in the art and are commercially available. Such machinestypically hold the metal tube between the open ends while a cuttinglaser, preferably under microprocessor control, cuts the pattern. Thepattern dimensions and styles, laser positioning requirements, and otherinformation are programmed into a microprocessor which controls allaspects of the process. After the stent pattern is cut, the stent istreated and polished using any number of methods well known to thoseskilled in the art. Lastly, the stent is then cooled until it iscompletely martensitic, crimped down to its un-expanded diameter andthen loaded into the sheath of the delivery apparatus.

It is believed that many of the advantages of the present invention canbe better understood through a brief description of a delivery apparatusfor the stent, as shown in FIGS. 1 and 2. FIGS. 1 and 2 show aself-expanding stent delivery apparatus I for a stent made in accordancewith the present invention. Apparatus I comprises inner and outercoaxial tubes. The inner tube is called the shaft 10 and the outer tubeis called the sheath 40. Shaft 10 has proximal and distal ends 12 and 14respectively. The distal end 14 of the shaft terminates at a luer lockhub 5. Preferably, shaft 10 has a proximal portion 16 which is made froma relatively stiff material such as stainless steel, Nitinol, or anyother suitable material, and an distal portion 18 which is made from apolyethylene, polyimide, pellethane, Pebax, Vestamid, Cristamid,Grillamid or any other suitable material known to those of ordinaryskill in the art. The two portions are joined together by any number ofmeans known to those of ordinary skill in the art. The stainless steelproximal end gives the shaft the necessary rigidity or stiffness itneeds to effectively push out the stent, while the polymeric distalportion provides the necessary flexibility to navigate tortuous vessels.

The distal portion 18 of the shaft has a distal tip 20 attached thereto.The distal tip 20 has a proximal end 34 whose diameter is substantiallythe same as the outer diameter of the sheath 40. The distal tip tapersto a smaller diameter from its proximal end to its distal end, whereinthe distal end 36 of the distal tip has a diameter smaller than theinner diameter of the sheath. Also attached to distal portion 18 ofshaft 10 is a stop 22 which is proximal to the distal tip 20. Stop 22can be made from any number of materials known in the art, includingstainless steel, and is even more preferably made from a highlyradiopaque material such as platinum, gold tantalum. The diameter ofstop 22 is substantially the same as the inner diameter of sheath 40,and would actually make frictional contact with the inner surface of thesheath. Stop 22 helps to push the stent out of the sheath duringdeployment, and helps the stent from migrating proximally into thesheath 40.

A stent bed 24 is defined as being that portion of the shaft between thedistal tip 20 and the stop 22. The stent bed 24 and the stent 50 arecoaxial so that the portion of shaft 18 comprising the stent bed 24 islocated within the lumen of the stent 50. However, the stent bed 24 doesnot make any contact with stent 50 itself. Lastly, shaft 10 has aguidewire lumen 28 extending along its length from its proximal end 12and exiting through its distal tip 20. This allows the shaft 10 toreceive a guidewire much in the same way that an ordinary balloonangioplasty catheter receives a guidewire. Such guidewires are wellknown in art and help guide catheters and other medical devices throughthe vasculature of the body.

Sheath 40 is preferably a polymeric catheter and has a proximal end 42terminating at a hub 52. Sheath 40 also has a distal end 44 whichterminates at the proximal end 34 of distal tip 20 of the shaft 18, whenthe stent is in its fully un-deployed position as shown in the figures.The distal end 44 of sheath 40 includes a radiopaque marker band 46disposed along its outer surface. As will be explained below, the stentis fully deployed when the marker band 46 is lined up with radiopaquestop 22, thus indicating to the physician that it is now safe to removethe apparatus I from the body. Sheath 40 preferably comprises an outerpolymeric layer and an inner polymeric layer. Positioned between outerand inner layers a braided reinforcing layer. Braided reinforcing layeris preferably made from stainless steel. The use of braided reinforcinglayers in other types of medical devices can be found in U.S. Pat. Nos.3,585,707 issued to Stevens on Jun. 22, 1971, 5,045,072 issued toCastillo et al. on Sep. 3, 1991, and U.S. Pat. No. 5,254,107 issued toSoltesz on Oct. 19, 1993, all of which are hereby incorporated herein byreference.

FIGS. 1 and 2 show the stent 50 as being in its fully un-deployedposition. This is the position the stent is in when the apparatus 1 isinserted into the vasculature and its distal end is navigated to atarget site. Stent 50 is disposed around stent bed 24 and at the distalend 44 of sheath 40. The distal tip 20 of the shaft 10 is distal to thedistal end 44 of the sheath 40, and the proximal end 12 of the shaft 10is proximal to the proximal end 42 of the sheath 40. The stent 50 is ina compressed state and makes frictional contact with the inner surface48 of the sheath 40.

When being inserted into a patient, sheath 40 and shaft 10 are lockedtogether at their proximal ends by a Touhy Borst valve 8. This preventsany sliding movement between the shaft and sheath which could result ina premature deployment or partial deployment of the stent. When thestent 50 reaches its target site and is ready for deployment, the TouhyBorst valve 8 is opened so that that the sheath 40 and shaft 10 are nolonger locked together.

The method under which apparatus 1 deploys stent 50 should be readilyapparent. The apparatus 1 is first inserted into a vessel so that thestent bed 24 is at a target diseased site. Once this has occurred thephysician would open the Touhy Borst valve 8. The physician would thengrasp the proximal end 12 of shaft 10 so as to hold it in place.Thereafter, the physician would grasp the proximal end 42 of sheath 40and slide it proximal, relative to the shaft 40. Stop 22 prevents thestent 50 from sliding back with the sheath 40, so that as the sheath 40is moved back, the stent 50 is pushed out of the distal end 44 of thesheath 40. Stent deployment is complete when the radiopaque band 46 onthe sheath 40 is proximal to radiopaque stop 22. The apparatus I can nowbe withdrawn through stent 50 and removed from the patient.

Although particular embodiments of the present invention have been shownand described, modification may be made to the device and/or methodwithout departing from the spirit and scope of the present invention.The terms used in describing the invention are used in their descriptivesense and not as terms of limitations.

1. A stent for insertion into a vessel of a patient, said stentcomprising: a tubular member with front and back open ends and alongitudinal axis extending there between, said member having a firstsmaller diameter for insertion into said vessel, and a second largerdiameter for deployment into said vessel, said tubular member comprisinga plurality of adjacent hoops each having proximal and distal open endsrelative to the longitudinal axis, said hoops comprising a plurality oflongitudinal struts and a plurality of loops connectingcircumferentially adjacent struts so as to form a series of peaks andvalleys in a substantially S or Z shaped pattern; and a plurality ofbridges connected between loops on adjacent hoops in a repeatingpattern, each said connection between the loops and hoops forming aconnection region, wherein the repeating pattern is a ratio of the totalnumber of loops for a given end of one of the adjacent hoops to thenumber of connection regions for the same end of the same hoop, andwherein the ratio is a whole number.
 2. The stent according to claim 1wherein the ratio is 2:1.
 3. The stent according to claim 1 wherein theratio is 4:1.
 4. The stent according to claim 1 wherein the ratio is8:1.
 5. The stent according to claim 1 wherein each bridge comprises anelongated linear strut member directly connected on each end to a firstend of a curved bridge loop member, the second end of each curved bridgeloop member being directly connected to the loop member at theconnection region.
 6. The stent according to claim 5 wherein the ratioof the circumference of the tubular member to the length of theelongated linear strut member is greater than
 4. 7. The stent accordingto claim 1 wherein the loops on adjacent hoops are oriented in the samedirection relative to the longitudinal axis.
 8. The stent according toclaim 1 wherein the loops on adjacent hoops are rotationally off-setrelative to the longitudinal axis.
 9. The stent according to claim 1wherein the whole number is an even number.
 10. The stent according toclaim 1 wherein the bridges are evenly spaced along the circumference ofthe tubular member.
 11. The stent according to claim 1 wherein the ratioof the circumference of the tube to the distance between the adjacenthoops is between 20:1 and 50:1.
 12. The stent according to claim 1wherein said stent is made from a superelastic alloy.
 13. The stentaccording to claim 12 wherein said alloy comprises from about 50.5percent to about 60 percent Nickel and the remainder comprisingTitanium.
 14. The stent according to claim 5 wherein said elongatedlinear strut member is tapered with the narrowest point at the centerpoint along its length.